Implantable stimulation devices are devices that generate and deliver electrical stimuli to body nerves and tissues for the therapy of various biological disorders, such as pacemakers to treat cardiac arrhythmia, defibrillators to treat cardiac fibrillation, cochlear stimulators to treat deafness, retinal stimulators to treat blindness, muscle stimulators to produce coordinated limb movement, spinal cord stimulators to treat chronic pain, cortical and deep brain stimulators to treat motor and psychological disorders, and other neural stimulators to treat urinary incontinence, sleep apnea, shoulder sublaxation, etc. The description that follows will generally focus on the use of the invention within a Spinal Cord Stimulation (SCS) system, such as that disclosed in U.S. Pat. No. 6,516,227. However, the present invention may find applicability in any implantable medical device system.
As shown in FIGS. 1A-1C, a SCS system typically includes an Implantable Pulse Generator (IPG) 100, which includes a biocompatible device case 30 formed of a conductive material such as titanium for example. The case 30 typically holds the circuitry and battery 26 necessary for the IPG to function, although IPGs can also be powered via external RF energy and without a battery. The IPG 100 is coupled to electrodes 106 via one or more electrode leads (two such leads 102 and 104 are shown), such that the electrodes 106 form an electrode array 110. The electrodes 106 are carried on a flexible body 108, which also houses the individual signal wires 112 and 114 coupled to each electrode. In the illustrated embodiment, there are eight electrodes on lead 102, labeled E1-E8, and eight electrodes on lead 104, labeled E9-E16, although the number of leads and electrodes is application specific and therefore can vary. The leads 102, 104 couple to the IPG 100 using lead connectors 38a and 38b, which are fixed in a non-conductive header material 36, which can comprise an epoxy for example.
As shown in the cross-section of FIG. 1C, the IPG 100 typically includes an electronic substrate assembly including a printed circuit board (PCB) 16, along with various electronic components 20 mounted to the PCB 16, some of which are discussed subsequently. Two coils (more generally, antennas) are generally present in the IPG 100: a telemetry coil 13 used to transmit/receive data to/from an external controller (not shown); and a charging coil 18 for charging or recharging the IPG's battery 26 using an external charger 50 (discussed further below). In this example, the telemetry coil 13 and charging coil 18 are within the case 30, as disclosed in U.S. Patent Publication 2011/0112610. (FIG. 1B shows the IPG 100 with the case 30 removed to ease the viewing of the two coils 13 and 18). However, the telemetry coil 13 may also be mounted within the header 36 of the IPG 100 (not shown).
FIG. 2 shows the IPG 100 in communication with external charger 50 just mentioned. The external charger 50 is used to wirelessly convey power to the IPG 100, which power can be used to recharge the IPG's battery 26. The transfer of power from the external charger 50 is enabled by a coil (antenna) 17. The external charger 50, like the IPG 100, also contains a PCB 70 on which electronic components 72 are placed. Again, some of these electronic components 72 are discussed subsequently. A user interface 74, including touchable buttons and perhaps a display and a speaker, allows a patient or clinician to operate the external charger 50. A battery 76 provides power for the external charger 50, which battery 76 may itself be rechargeable. The external charger 50 can also receive AC power from a wall plug. A hand-holdable case 77 sized to fit a user's hand contains all of the components.
Power transmission from the external charger 50 to the IPG 100 occurs wirelessly, and transcutaneously through a patient's tissue 25, via inductive coupling. FIG. 3 shows details of the circuitry used to implement such functionality. Coil 17 in the external charger 50 is energized via charging circuit 122 with a constant non-data-modulated AC current, Icharge, to create an AC magnetic charging field. This magnetic field induces a current in the charging coil 18 within the IPG 100, which current is rectified (132) to DC levels, and used to recharge the battery 26, perhaps via a charging and battery protection circuit 134 as shown. The frequency of the magnetic charging field can be perhaps 80 kHz or so. When charging the battery 26 in this manner, is it typical that the case 77 of the external charger 50 touches the patient's tissue 25, although this is not strictly necessary.
The IPG 100 can also communicate data back to the external charger 50 during charging using reflected impedance modulation, which is sometimes known in the art as Load Shift Keying (LSK). Such back telemetry from the IPG 100 can provide useful data concerning charging to the external charger 50, such as the capacity of the battery 26, or whether charging is complete and the external charger 50 can cease.
Control circuitry 140 in the IPG 100 monitors the battery voltage, Vbat, and with the assistance of LSK module 155, produces LSK data. The control circuitry 140 can include a microcontroller for example, and may be associated with Analog-to-Digital (A/D) conversion circuitry to process and interpret the battery voltage. LSK module 155 preferably operates as software in the control circuitry 140, and assesses the incoming battery voltage to produce appropriate LSK data at appropriate times. Such LSK data is sent as a serial string of bits along line 99 to the gates of load transistors 141 and 142. The LSK data modulates the state of transistors 141 and 142, which in turn modulates the impedance of the coil 18. When LSK data=1, the transistors 141 and 142 are on (shorted) which shorts each end of the coil 18 to ground. When LSK data=0, the transistors are off (opened). The impedance of the coil 18 may also be modulated by a single transistor in series with the coil 18, which modulates the impedance by opening the coil, as shown in dotted lines.
Such modulation of the charging coil 18 is detectable at the external charger 50. Due to the mutual inductance between the coils 17 and 18, any change in the impedance of coil 18 affects the voltage needed at coil 17, Vcoil, to drive the charging current, Icharge: if coil 18 is shorted (LSK data=1), Vcoil increases to maintain Icharge; if not shorted (LSK data=0), Vcoil decreases. In this sense, the impedance modulation of coil 18 is “reflected” back to the transmitting coil 17, and thus data can be said to be “transmitted” from the IPG 100 to the external charger 50, even if not transmitted in the traditional sense. An example Vcoil waveform arising from transmission of an example sequence (LSK data=01010) is shown at the bottom of FIG. 3, and shows the data states as modulated by the ˜80 kHz frequency of the magnetic field.
The Vcoil waveform is processed at demodulation circuitry 123 to recover the transmitted LSK data. To be reliably detected, the difference in coil voltage (ΔV) between the transmitted ‘0’ (Vcoil0) and ‘1’ (Vcoil1) states must as a practical matter be greater than a threshold voltage inherent in the demodulator 123, Vt1. Depending on the particularly of the circuitry, Vt1 can be rather small, ranging from 50 mV to 100 mV for instance, and can be statistically determined based on suitable bit error rates for LSK transmission.
The serial stream of demodulated bits is then received at control circuitry 144 operating in the external charger 50, so that appropriate action can be taken. The control circuitry 144 can again include a microcontroller for example. For example, if an alternating stream of bits is received (01010101 . . . ), this might be interpreted by the control circuitry 144 that the battery 26 in the IPG 100 is full, and therefore that charging can cease. In such an instance, the control circuitry 144 can suspend the production of the magnetic charging field (i.e., setting Icharge to 0), and may notify the user of that fact (by a graphical display, an audible beep, or other indicator).
Because LSK telemetry works on a principle of reflection, LSK data can only be communicated from the IPG 100 to the external charger 50 during periods when the external charger is active and is producing a magnetic charging field.
An issue arising when inductive coupling is used for power transmission relates to the coupling between the coils 17 and 18 in external charger 50 and the IPG 100. Coupling, generally speaking, comprises the extent to which power expended at the transmitting coil 17 in the external charger 50 is received at the coil 18 in the IPG 100. It is generally desired that the coupling between coils 17 and 18 be as high as possible: higher coupling results in faster charging of the IPG battery 26 with the least expenditure of power in the external charger 50. Poor coupling is disfavored, as this will require high power drain (i.e., a high Icharge) in the external charger 50 to adequately charge the IPG battery 26. The use of high power depletes the batteries 76 (if any) in the external charger 50, and more importantly can cause the external charger 50 to heat up, and possibly burn or injure the patient.
Coupling depends on many variables, such as the permeability of the materials used in the external charger 50 and the IPG 100, as well materials inherent in the environment. Coupling is also affected by the relative positions of the external charger 50 and IPG 100, as shown in FIGS. 4A-4D. For best coupling, it is preferred that axes around which coils 17 and 18 are wound (17′ and 18′) are parallel and collinear, and that the coils 17 and 18 as close as possible (d1) to each other, as shown in FIG. 4A. Distance d1 indicates the depth between the external charger 50 and the IPG 100, and is generally constant given that the external charger is generally placed on the patient's tissue 25, and that the IPG 100 has been implanted at a particular depth. Deviations from these ideal conditions will generally reduce coupling, as shown in FIGS. 4B-4D. In FIG. 4B for instance, the coil axes 17′ and 18′ are not collinear, but instead are laterally offset (x). In FIG. 4C, the coil axes 17′ and 18′ are not parallel, but instead have an angle θ between them. In FIG. 4D, the coil axes 17′ and 18 are parallel and collinear, but the IPG 100 is relatively deep (d2).
In any of these non-ideal cases 4B-4D, coupling will be reduced, meaning that the external charger 50 must output more power (e.g., Icharge must be higher) to affect the same charging rate of the IPG's battery 26. In a closed loop charging system, the relative degree of coupling between the external charger 50 and the IPG 100 is assessed, and is used to change the output power accordingly. Prior systems have quantified this coupling in different ways. In one approach, data indicative of the coupling is read at the IPG 100 and telemetered back to the external charger 50, which again can adjust its output power accordingly. See U.S. Patent Publication 2011/0087307. But this approach adds additional complexity to the system. Applicants have found a new way of quantifying coupling in the external charger/IPG system that doesn't rely on telemetry of coupling parameters from the IPG, which is easy to implement, and which therefore results in an improved closed loop charging system.